Active matrix X-ray imaging array

ABSTRACT

A digital detector for radiography and fluoroscopy is disclosed. The detector includes a large area, flat panel that easily fits into a conventional X-ray room bucky tray. The detector utilizes a layer of photoconductor (i.e. a-Se in the preferred embodiment) to detect X-rays and convert the X-ray energy to charge, and an active matrix TFT array in the form of a very large area integrated circuit, for readout of the charge. A dual gate structure is used for the TFT array wherein the top gate is formed as an extension of the pixel electrode, so as to provide high voltage protection of the TFT. An integrated pixel storage capacitor is provided for enhanced absorption of X-ray energy with low pixel voltage, low leakage current and a large charge leakage time constant. In a preferred embodiment, the integrated pixel storage capacitor is created by overlapping the pixel the pixel electrode with an adjacent gate line or a separate ground line of the active matrix readout array. Image charge collection efficiency is enhanced by manipulating the electric field distribution in the photoconductor layer so that image charges land on the pixel electrodes, and not on the TFT readout devices.

FIELD OF THE INVENTION

This invention relates in general to medical diagnostic imaging systems,and more particularly to a selenium active matrix universal readoutarray imager.

BACKGROUND OF THE INVENTION

Despite the development of recent medical imaging modalities, such ascomputed tomography (CT), ultrasound, nuclear medicine and magneticresonance imaging (MRI), all of which are digital, X-ray imaging systemsremain an important tool for medical diagnosis. Although the majority ofX-ray imaging systems in current use are of analog design, digitalradiology is an area of considerable recent growth. Digital radiologyprovides significant advantages over its analog counter-part, such as:easy comparison of radiological images with those obtained from otherimaging modalities; the ability to provide image networking within ahospital for remote access and archiving; facilitating computer aideddiagnosis by radiologists; and facilitating teleradiology (ie. remotediagnostic service to poorly populated regions from a central facility).

There are currently two commercial approaches to digitalradiography--(1) the digitization of a signal from a video cameraoptically coupled to a an X-ray imaging intensifier, and (2) stimulablephosphor systems. Prior art intensifier systems permit instant readoutwhereas prior art stimulable phosphor systems require the operator tocarry a cassette to a reader. Neither of these systems provide imagequality which is acceptable for all applications.

Digital systems based on the use of X-ray image intensifiers suffer fromthe following disadvantages: the bulky nature of the intensifier oftenimpedes the clinician by limiting access to the patient and prevents theacquisition of important radiographic views; loss of image contrast dueto X-ray and light scattering (i.e. veiling glare); and geometric (pincushion) distortion on the image due principally to the curved inputphosphor.

Another prior art X-ray imaging modality which is currently experiencingrenewed interest, is the use of amorphous selenium photoconductors as analternative to phosphors. Xeroradiography, (i.e. the use of amorphousselenium (a-Se) plates which are read out with toner), was a technicaland commercial success in the early 1970's. Xeroradiography is no longercommercially competitive. This is believed to be because of the tonerreadout method, and not because of the underlying properties of a-Se.Commercial as well as scientific interest in a-Se has recently revived.For example, Philips has announced the commercial availability of ana-Se drum scanner for chest radiography based on earlier work at itsresearch laboratories in Aachen. Kodak uses an a-Se plate readout with aphosphor coated toner and laser scanner for the preparation of highlydetailed mammography images which are free from significant artifacts.3M have also published preliminary descriptions of their work on laserdischarge readout of a-Se. This work is related to much earlierpublications by (1) Korn et al, "A method of electronic readout ofelectrophotographic and electroradiographic images", Journal of AppliedPhotographic Engineering, 4, 178-182 (1978); (2) Zermeno et al "Laserreadout of electrostatic images", In: Application of OpticalInstrumentation to Medicine VII, Edited by J. Gray, et al, SPIE 173,81-87 (1979); and (3) DeMonts et al, "A new photoconductor imagingsystem for digital radiography", Medical Physics, 16, 105-109 (1989).

The basis of all existing medical X-ray imaging systems is a phosphorlayer or "screen". X-rays absorbed by the screen release light whichmust reach the surface to create an image. The lateral spread of lightis limited only by diffusion and hence is related to the thickness ofthe screen. Thus, the thicker the screen (which is desirable to increasethe quantum absorption efficiency), the more blurry the image will be.This represents a loss of high frequency image information in prior artphosphor systems which is fundamental and largely irreversible. Thisloss can be alleviated to some extent by using a phosphor such as CsIwhich can be grown in the form of a fibre optic.

A better method has been discovered for eliminating blurring, whichinvolves using a structureless photoconductor to detect X-rays. X-raysinteracting in the photoconductor release electron-hole pairs which aredrawn directly to the surfaces of the photoconductor by an appliedelectric field. The latent charge image on the photoconductor surface istherefore not blurred significantly even if the photoconductor layer ismade thick enough to absorb most incident X-rays. Amorphous selenium(a-Se) is the most highly developed photoconductor for X-rayapplications. Its amorphous state maintains uniform characteristics tovery fine levels over large areas. A large area detector is essential inradiography since no means are provided to focus the X-rays, therebynecessitating a shadow X-ray image which is larger than the body part tobe imaged.

One area of intense research in the field of photoconductor X-raydetectors, is the development of systems for charge readout. Antonuk etal disclosed the concept of an X-ray imaging detector which utilizesactive matrix arrays for charge readout, as described in the followingpublications: (1) "Signal, noise, and readout considerations in thedevelopment of amorphous silicon photodiode arrays for radiotherapy anddiagnostic imaging", Medical Imaging V: Imaging Physics, SPIE 1443,108-119 (1991), (2) "High resolution, high frame rate, flat panel TFTarrays for digital X-ray imaging", Medical Imaging 1994: Physics ofMedical Imaging, Rodney Shaw, Editor, Proceedings of SPIE, 2163, 118-128(1994) and (3) "Demonstration of megavoltage and diagnostic X-rayimaging with hydrogenated amorphous silicon arrays", Medical Physics 19,1455-1466 (1992). Their initial research has subsequently been developedby others: Ichiro Fujieda, Robert A. Street, Richard L. Weisfield, SteveNelson, Per Nylen, Victor Perez-Mendez and Gyuseong Cho, "Highsensitivity readout of 2d a-Si image sensors", Jpn. J. Appl. Phys. 32,198-204 (1993); Henri Rougeot, "Direct X-ray photoconversion processes",In: Digital imaging: AAPM 1993 Summer School Proceedings Ed: WilliamHendee and Jon Trueblood (AAPM monograph 22, Medical Physics Publishing,1993) pp. 49-96; UW Schiebel, N Conrads, N Jung, M Weilbrecht, HWieczorek, T T Zaengel, M J Powell, I D French and C Glasse"Fluoroscopic X-ray imaging with amorphous silicon thin-film arrays",Medical Imaging 1994: Physics of Medical Imaging, Rodney Shaw, Editor,Proc. SPIE, 2163, 129-140 (1994); and M J Powell, I D French, J RHughes, N C Bird, O S Davies C Glasse and J E Curran, "Amorphous siliconimage sensor arrays", Mat. Res. Soc. Symp. Proc. 258, 1127-1137 (1992).

In these prior art systems a phosphor screen (preferably a structuredCsI layer) is used to absorb X-rays, and the resultant light photons aredetected by an active matrix array with a single photodiode andtransistor at each pixel. Antonuk coined the acronym MASDA for"Multi-element Amorphous Silicon Detector Array".

SUMMARY OF THE INVENTION

According to the present invention, a digital detector is provided whichperforms all of the currently available radiological modalities,radiography (including rapid sequence radiography) and fluoroscopy. Thedetector comprises a large area, flat panel which easily fits into theconventional X-ray room bucky tray. The detector utilizes a layer ofphotoconductor (ie. a-Se in the preferred embodiment) to detect X-raysand convert the X-ray energy to charge, and an active matrix TFT arrayin the form of a very large area integrated circuit, for readout of thecharge. The broad concepts which led to this invention are disclosed inthe following article: W. Zhao and J. A. Rowlands, "Digital RadiologyUsing Self-Scanned Readout of Amorphous Selenium", in Medical ImagingVII: Physics of Medical Imaging, SPIE 1896, 114-120 (1993). However,certain inventive aspects of implementation of the device are notdisclosed in this prior article and form the basis of the presentapplication.

According to one aspect of the present invention, a dual gate structureis utilized for providing high voltage protection of the TFTs. Theadditional gate is formed as an extension of the pixel electrode, andoverlies a predetermined thickness of dielectric over the semiconductorchannel. When excessive charge is collected by the electrode, the TFTturns ON so that a high leakage current drains away the excess charge onthe pixel electrode.

According to a further aspect of the invention, an integrated pixelstorage capacitor is provided for enhanced absorption of X-ray energywith low pixel voltage, low leakage current and hence a large chargeleakage time constant. In the preferred embodiment, the integrated pixelstorage capacitor is created by overlapping the pixel electrode with anadjacent gate line or a separate ground line of the active matrixreadout array.

According to another aspect of the invention, image charge collectionefficiency is improved by manipulating the electric field distributionin the photoconductor layer so that image charges land on the pixelelectrodes, and not on the TFT readout devices.

According to yet another aspect of the invention, a photo-timer isintegrated into the imaging detector for measuring X-ray exposure.

As discussed in greater detail below, because an electrostatic X-rayimage transducer is utilized, the system of the preferred embodimentprovides higher resolution images than phosphor based systems, eventhose using structured CsI. The signal-to-noise ratio of the prior artMASDA system and the system of the preferred embodiment are essentiallyidentical since the X-ray-to-charge conversion gain is the same for both(assuming CsI and a-Si:H for MASDA and a-Se for the system of thepreferred embodiment). Thus, the overall image quality of the systemaccording to the present invention is believed to be considerably betterthan that produced using the prior MASDA device.

Furthermore, the requirements for manufacture of the system of thepreferred embodiment are favourable when compared to the prior art MASDAsystem. Firstly, MASDA requires a CsI structure which is more difficultin principle to manufacture than a uniform layer of a-Se. Secondly,because X-rays are converted directly to electrons by a-Se, the need forphotodiodes at each pixel is eliminated and the active matrix array canbe simplified. This leads to further simplifications in the system ofthe present invention, as compared to the prior art MASDA device,thereby resulting in more economical manufacturing.

These and other aspects of the invention are described in greater detailbelow.

BRIEF DESCRIPTION OF THE DRAWINGS

A detailed description of the preferred embodiment is provided belowwith reference to the following drawings, in which:

FIG. 1A is a schematic plan view of the imaging arraying to thepreferred embodiment;

FIG. 1B is an equivalent circuit for the imaging array of FIG. 1A;

FIG. 2 is cross sectional view through a single pixel of the array shownin FIG. 1;

FIG. 3 shows the I-V characteristics of the high voltage protected TFTaccording to the present invention;

FIG. 4A is a cross sectional view through two pixels of the array shownin FIG. 1, illustrating improved fill factor by bending electric fieldlines using guard rails, in accordance with the preferred embodiment;

FIG. 4B is a plan view of the top layer of the array showing thedisposition of the guard rails;

FIG. 4C shows an alternative embodiment in which fill factor is improvedby the bending of electric field lines using the charge trappingproperties of the top dielectric material between pixel electrodes;

FIG. 5 is a plan view of an arrangement of bias electrode for biasing aphotoconductor layer of the preferred embodiment and providing dosemeasurements in accordance with an alternative embodiment;

FIGS. 6A and 6B are two alternative cross-sectional views through theline VI--VI in FIG. 5; and

FIG. 7 is a schematic of a photo-timer and circuit arrangement fordose/dose rate measurement, according to the embodiment of FIGS. 5 and6.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

With reference to FIG. 1, an active matrix 10 is shown comprising aplurality of pixels, each comprising a pixel electrode 12, storagecapacitor 14 and thin film transistor (TFT) 16. An external scanningcontrol circuit 18 turns on the TFTs 16 one row at a time via aplurality of control lines 19, for transferring the image charge fromthe pixels to a plurality of data lines 20, and then to respectiveexternal charge amplifiers 22. At the same time, the input (virtualground) of the charge amplifiers 22 resets the potential at each pixelelectrode 12. The resulting amplified signal for each row is multiplexedby a parallel-to-serial converter or multiplexer 24, and thentransmitted to an analog-to-digital converter or digitizer 26.

Each TFT 16 comprises 3 electrical connections: the drain (D) isconnected to the pixel electrode 12 and pixel storage capacitor 14; thesource (S) is connected to a common data line 20 shared by all TFTs ofthe same column, and also to an external charge sensitive amplifier 22;and the gate (G) is used for control of the "on" and "off" state of theTFT 16. Usually, 10 V and -5 V is applied to turn on and off the TFT 16respectively.

The scanning control circuit 18 may be fabricated as a single crystalsilicon integrated circuit which is wire bonded to the active matrix TFTarray. The charge amplifiers 22 and multiplexer 24 may also befabricated as a single crystal silicon integrated circuit which is wirebonded to the active matrix array.

Turning now to FIG. 2, the structure of a single pixel is shown of thelarge area integrated circuit active matrix.

First, a metal layer (preferably Cr or Al) is deposited (by thermalevaporation or sputtering) on a glass substrate 28 and patterned usingphotolithography to form the gate regions (G) for the array of TFTS. Asdiscussed in greater detail below, the gate line of an adjacent pixelmay be extended so that the gate line and the pixel electrode 12 form anintegrated pixel storage capacitor 14 with insulating layer 30 extendingtherebetween. Alternatively, separate ground return electrodes forstorage capacitor may be formed between gate electrode lines on thefirst metal layer. The insulating layer 30 is deposited using PECVD(Plasma Enhanced Chemical Vapour Deposition) or thermal evaporation. Theinsulating material can be SiO₂, Si₃ N₄, or alternate layers of both.The thickness of the layer is typically 0.1-0.5 μm.

Next, the drain (D) and source (S) metal layers are deposited (bythermal evaporation or sputtering) and patterned using photolithographyto form drain and source contact pads for the TFT, the pixel electrodesand source (i.e. data) lines. The preferred material for the D and Scontact pads is Cr, and an extra coating of Al is preferably added tothe source lines to reduce the source line resistance. Next, asemiconductor layer 32, being several hundred angstroms thick, isdeposited (e.g. using thermal evaporation or sputtering in the case ofCdSe) and then patterned using photolithography to form the TFT channel(e.g. 30 μm wide and 50 μm long, although the illustrated TFT geometryrepresents only one possible embodiment of the invention).

The above-described deposition procedure is used for the drain, andsource metal and semiconductor fabrication steps for a bottom D and Scontact TFT structure. The two deposition steps can be reversed to forma top contact structure.

Next a dielectric layer 34 (SiO₂, Si₃ N₄ or alternate layers of both) isdeposited (using PECVD or thermal evaporation) with a thickness of 0.3-5μm. Then, the dielectric on top of the pixel electrode is etched away toexpose the pixel electrode.

The final top metal layer (preferably Al, or ITO) of the TFT isdeposited using sputtering or thermal evaporation, and patterned usingphotolithography to form the pixel electrode 12 (which is the bottompixel electrode since the dielectric in this region has been etchedaway). As discussed in greater detail below, according to the preferredembodiment, the pixel electrode 12 extends over the top gate dielectriclayer 34 so as to form a dual gate TFT structure. A blocking layer maybe formed by thermal oxidization the top metal (Al) layer for preventingnegative charge injection from the pixel electrode to the X-rayphotoconductor.

A uniform layer of X-ray sensitive photoconductor 36 is then directlydeposited on the surface of the active matrix by thermal evaporation, toa thickness of approximately 500 μm. Preferably, the photoconductor isfabricated from amorphous selenium (a-Se).

A top bias electrode 38 is deposited (e.g. by thermal evaporation) ontothe photoconductor layer 36 with appropriate blocking contact so thatcharge generated in the bulk of the photoconductor can flow to the biaselectrode, with no charge injection from the bias electrode into thephotoconductor. Several types of metal may form the blocking contactwith selenium, such as Au, Indium, etc. An alternative embodiment is todeposit a thin layer (several hundred angstroms) of insulator (e.g.CeO₂) on the surface of the selenium before the bias electrode isdeposited, wherein the thin insulating layer serves as a blocking layer.

Returning briefly to FIG. 1B, the selenium layer 36 and top biaselectrode 38 are shown schematically as a photodiode connected to a highbias voltage (HV) at the cathode of each pixel.

During X-ray irradiation, the X-ray energy is absorbed by the X-rayphotoconductor 36 and electron-hole pairs are created. Under the appliedelectric field created by the difference in potential between biaselectrode 38 and pixel electrode 12, the radiation generated charges aredrawn to the surfaces of the photoconductor 36 and collected on pixelelectrode 12. The difference in charge at each pixel represents theX-ray image.

As discussed above, the pixel electrode is connected to the drain (D) ofthe TFT 16. During each readout, the potential of the pixel electrode isreset, through the TFT, to a ground potential by the virtual groundinput of the charge amplifier 22.

For fluoroscopy applications, a high voltage is constantly applied tothe bias electrode 38 and the imaging detector is scanned in real time(i.e. 30 frames per second). The images are acquired continuously inevery 1/30 second frame and are processed and displayed in real time.

For radiography applications, a high voltage is applied to the biaselectrode 38 and the scanning is suspended (i.e. all TFTs 16 are turnedoff) during X-ray exposure. Scanning is resumed immediately after theexposure in order to readout the image.

For a-Se, the photoconductor layer 36 needs to be of a thickness in theorder of 500 μm in order to absorb most of the incident X-rays. Thus,the bias voltage applied to electrode 38 must be in the order of 5000volts under an electric field of 10 V/μm. Under abnormal conditions(e.g. a false prolonged X-ray exposure when all TFTs 16 are turned off),the potential on each pixel (V_(P)) can reach a damaging high value(e.g. 1000 volts).

The CdSe TFTs 16 of the preferred embodiment can maintain normalfunctions at V_(P) up to approximately 200 volts. Thus, it is necessaryto ensure that even under false, abnormal conditions, V_(P) does notexceed 100 volts.

As discussed briefly above, and as shown with reference to FIG. 2, adual gate structure is utilized to protect the TFT 16 from high voltagedamage. In particular, the pixel electrode 12 (which is connected to theTFT drain (D)), extends over the top of the TFT 16 and acts as a secondgate. The top gate voltage is equivalent to the pixel voltage (i.e.V_(TG) =V_(P)).

By adjusting the thickness of the top dielectric layer 34, the effect ofV_(P) on the transfer characteristics of the TFT 16, can be controlled.The top dielectric layer 34 is usually 5 to 10 times the thickness ofthe bottom gate dielectric layer for high voltage protection at a pixelpotential of 100 volts. FIG. 3 shows the I_(D) -V_(G) characteristiccurve for a dual gate TFT at different values of V_(P). Under normalimaging conditions (i.e. V_(P) <10 V), the bottom gate control pulsecauses the TFT 16 to turn on and off correctly. However, if V_(P)exceeds 100 volts, the bottom gate control pulse is no longer able toturn off TFT 16. In this case, the high leakage current drains away theexcess charge on the pixel electrode 12 and V_(P) never reaches adangerously high potential.

The relationship between maximum pixel voltage V_(P)(max) and dielectricthickness may be expressed as follows: ##EQU1## where ε_(it) is thedielectric constant of dielectric layer 34, d_(it) is the thickness ofdielectric layer 34, V_(p)(max) is the maximum voltage to be applied tothe pixel, ε_(ib) dielectric constant of the dielectric layer 30, andd_(ib) is the thickness thereof. When V_(G)(off) (usually -5 V) isapplied to the bottom gate (G), it is desired that the TFT willnonetheless turn on when the voltage applied to pixel electrode 12reaches V_(p)(max). V_(th) is a constant representing the minimumvoltage which when applied to the bottom gate (G) will turn on the TFTwhen V_(p) =0. Thus for a dielectric layer 30 having thickness in therange of 0.1 to 0.5 μm, a maximum pixel voltage of 100 volts, and theconstant V_(th) of 10 volts, the dielectric layer 34 will have athickness of 1-5 μm, given the same dielectric as the dielectric layer30.

Another consequence of making the photoconductor layer 36 thick toabsorb as much X-ray energy as possible, is that a small sensorcapacitance is created for each pixel (e.g. approximately 0.01 pF). Thiscan result in three problems. Firstly, the pixel voltage V_(P) on thedrain (D) of the TFT 16 rises rapidly with the image charge (e.g.approximately 100 V/pC) because of the small pixel capacitance (i.e. thesum of the sensor capacitance C_(Se) and the coupling capacitancebetween the gate and drain of the TFT (C_(GD))), which in turn can causehigh voltage damage to the TFTs 16 and the external electronics (e.g.scanning control circuit 18, charge amplifiers 22, multiplexer 24).Secondly, when each TFT 16 is turned off, charge injection to the pixelelectrode 12 by the negative edge of the gate pulse output from scanningcontrol circuit 18 (e.g. 15 volts), results in a negative potential onthe pixel and thus a small forward bias between the gate (G) and drain(D). This can cause a significant increase in the leakage current forthe TFT 16. Thirdly, the charge leakage time constant for each pixelC_(P) ×R_(off) (approximately 10¹³ Ω) is 100 mS. For radiographyapplications, the pixels that are read out last will thus experiencesignificant signal loss due to the short leakage time constant.

According to the preferred embodiment, an integrated pixel storagecapacitance (C_(ST)) is provided on the TFT active matrix array, byoverlapping the pixel electrode 12 with the gate line (G) of an adjacentpixel, as shown in FIG. 1 and in FIG. 2 on the left where storagecapacitor 14 is formed by overlapping pixel electrode 12 with anextension of the gate line (G) of an adjacent pixel. As an alternativeto overlapping the pixel electrode 12 with the adjacent gate line, aseparate ground line may be utilized. A large pixel capacitance resultsfrom the thin insulating layer 30 (typically 0.1-0.5 μm), resulting in astorage capacitance C_(ST) in the range of 0.5-1 pF, which is 20 timeslarger than C_(GD), and two orders of magnitude larger than thecapacitance of the photoconductor layer 36. The value of C_(ST) isachieved by extending the pattern of the gate electrode (or a separateground line), under the region of each pixel electrode 12 when the sizeof the pixel electrode is larger than 200 μm (e.g. for fluoroscopy andgeneral radiography). For mammography applications, since the pixel sizemust be smaller (in the order of 50 μm), thinning of the insulator isneeded in addition to extending the gate electrode.

The large integrated pixel storage capacitance C_(ST) ensures, firstly,that the pixel voltage V_(P) does not rise more than 2 V/pC with imagecharge, and thus does not reach a damagingly high potential underdiagnostic X-ray exposure levels. Secondly, the voltage on the pixelelectrodes returns to near ground potential after the TFTs 16 are turnedoff, thereby ensuring a low leakage current. Thirdly, the charge leakagetime constant is approximately 10 seconds, and thus does not cause anysignificant signal loss for radiography applications.

Turning to FIG. 4, a cross sectional view, is provided similar to FIG.2, through two adjacent pixels. However, the section of FIG. 2 extendsthrough storage capacitor 14, while the section of FIG. 4 does not.According to the embodiment illustrated in FIG. 4, a plurality ofparallel rails 40 are deposited as a grid adjacent the pixel electrodes12, so as to overlay the source lines (S). Image charge collectionefficiency in an active matrix sensor array, is controlled by the fillfactor (i.e. the fraction of the area of each pixel that is occupied bythe pixel electrode 12). The fill factor of a typical CdSe TFT array isapproximately 80% for a 200 μm square pixel. Most of the remainder ofeach pixel is occupied by the source lines (S). By applying a potentialon the grid 40 that is significantly higher than the pixel potential,the electric field distribution in the photoconductive layer 36 may bemanipulated so that image charges only land on the pixel electrodes 12,and not on the source lines (S). As seen in FIG. 4, the field lines 42may be caused to bend toward the pixel electrodes 12 and thus increasethe effective fill factor. In practice, the potential applied to thegrid 40 must be sufficient to cause a noticeable increase in chargecollection efficiency of the pixel electrode 12 (e.g. typically in theorder of several hundred volts).

A plan view of the grid 40 is shown in FIG. 4B.

With reference to FIG. 4C, instead of utilizing a grid to bend thefield, as in the embodiment of FIGS. 4A and 4B, the charge trappingproperties of the top dielectric material of the pixel electrodes may beutilized to bend the electric field. More particularly, after theconstruction of the detector is completed, a seasoning process isperformed. To perform this seasoning, the detector is exposed to largedoses of X-rays (or visible light if the top bias electrode 38 issemitransparent, e.g. Au), with the TFTs 16 all turned on and with anelectric field applied to the selenium photoconductor 36. The holescreated in the photoconductor 36 are drawn to the bottom surfacethereof, either landing on the pixel electrodes 12 or becoming trappedby the dielectric material 34 between pixel electrodes. Holes which landon the pixel electrodes 12 are drained away through the turned-on TFTs16 and the holes trapped at the insulator 34 generate a surfacepotential which increases with the number of holes trapped. When thepotential rises to a level wherein further holes are repelled from theinsulator, the system has reached equilibrium. Since the trapping ofholes is a long term effect, when the detector is used for imaging afterthis stage, X-ray created holes will prefer to land on the pixelelectrodes 12 and the effective fill-factor of the system is therebyincreased to nearly 100%. This seasoning process may be performed onceafter the detector is constructed, or may be performed at the beginningof each day during which imaging is expected to be performed. It isfurther contemplated that repeated seasoning may not be necessary afterlong term usage of the device since the dark current of selenium may beenough to perform hole repelling after a sufficient term of use.

According to a further aspect of the present invention, means formeasuring X-ray exposure dosages may be incorporated into the design ofthe active matrix flat panel detector so as to perform photo timingfunctions simultaneously with image detection.

FIG. 5 shows the top view of top sensor bias electrode 38 which, asdiscussed above, is connected to a high voltage power supply. Aplurality of smaller electrodes 42 (e.g. preferably 3 for chestradiography) provide regions of X-ray dose measurement.

The bias electrode 38 is connected to a DC high voltage (HV) powersupply. Each phototimer electrode 42 is connected to its own dose/doserate measurement circuit. As shown in FIG. 7, each electrode 42 isconnected to the inverting input of an amplifier 71 which is powered bya pair of isolated power supplies, for providing +15 V and -15 V withthe ground reference set at the DC HV bias potential applied to thephotoconductor 36. The inverting input of amplifier 71 is at the samepotential as its non-inverting input, which is connected to the DC HVbias. Therefore, electrode 42 is at the same potential as electrode 38.When X-rays are absorbed by the photoconductor 36, current generated inthe region of the phototimer flows to the amplifier 71 (since a closedloop circuit is provided by the storage capacitors 14 and C_(GD)). Inthe case of fluoroscopy, the X-ray generated current is measured with afeed-back resistor 73 at the amplifier, resulting in an output voltagesignal which is, in turn, measured by a circuit in the X-ray generator(not shown) to determine whether it is the expected value and thereforewhether to change the X-ray tube current. In the case or radiography,the photocurrent generated during a short pulse (a fraction of a second)of X-ray exposure is integrated by the feedback capacitor 75 of theamplifier. When the amplifier 71 output voltage (also monitored by acircuit in the X-ray generator) reaches a preset value (i.e.proportional to the preset X-ray exposure dosage), the X-ray generatorwill turn off the X-rays.

The imaging mode (fluoroscopy or radiography) is selected electronicallyby a relay 77. Since the relay 77 is connected to the amplifier circuit,it has to be operated by a control signal with the same reference (i.e.DC HV potential).

In the cross sectional view of the FIG. 6A embodiment, a gap 43 isprovided for isolating the phototimer bias electrodes 42 from the commontop bias electrode 38, whereas in the cross-sectional view of the FIG.6B embodiment no gap for electric field application is shown (whenviewed in plan), and the necessary isolation between electrodes isprovided by an additional insulation layer 44.

Other embodiments and variations of the invention are possible. All suchmodifications and variations are believed to be within the sphere andscope of the invention as defined by the claims appended hereto.

The embodiments of the invention in which an exclusive property of privilege is claimed are defined as follows:
 1. An active matrix imager, comprising:a) an array of thin film transistors disposed in a plurality of rows and columns, each of said transistors having a control terminal and a pair of signal terminals; b) a dielectric layer overlying each of said thin film transistors; c) scanning control circuit means having a plurality of control lines, respective ones of said control lines being connected to the control terminals of each of the thin film transistors in respective ones of said rows; d) read out circuit means having a plurality of data lines, respective ones of said data lines being connected to a first one of said pair of signal terminals of each of the thin film transistors in respective ones of said columns; e) a plurality of pixel electrodes respectively connected to a second one of said pair of signal terminals of each of the thin film transistors in said array of thin film transistors; f) a plurality of storage capacitors connected to respective ones of said pixel electrodes; g) a photoconductive layer overlying said plurality of pixel electrodes and said dielectric layer, wherein electron-hole pairs are created in response to exposing said photoconductive layer to radiation; h) a bias electrode overlying said photoconductive layer; i) first voltage means for establishing a high voltage difference between said bias electrode and respective ones of said pixel electrodes, whereby charges created by said electron-hole pairs are collected on respective ones of said pixel electrodes and stored on respective ones of said storage capacitors, the amount of said collected charges being proportional to intensity of said radiation exposure; and j) means overlying said first one of said pair of signal terminals of each of the thin film transistors in respective ones of said columns for establishing an electric field for repelling said charges in the vicinity of said first one of said pair of signal terminals toward said pixel electrodes, wherein said means overlying said first one of said pair of signal terminals of each of the thin film transistors further comprises a plurality of grid lines connected to a source of opposite polarity voltage to said bias electrode.
 2. An active matrix imager, comprising:a) an array of thin film transistors disposed in a plurality of rows and columns, each of said transistors having a control terminal and a pair of signal terminals; b) a dielectric layer overlying each of said thin film transistors; c) scanning control circuit means having a plurality of control lines, respective ones of said control lines being connected to the control terminals of each of the thin film transistors in respective ones of said rows; d) read out circuit means having a plurality of data lines, respective ones of said data lines being connected to a first one of said pair of signal terminals of each of the thin film transistors in respective ones of said columns; e) a plurality of pixel electrodes respectively connected to a second one of said pair of signal terminals of each of the thin film transistors in said array of thin film transistors; f) a plurality of storage capacitors connected to respective ones of said pixel electrodes; g) a photoconductive layer overlying said plurality of pixel electrodes and said dielectric layer, wherein electron-hole pairs are created in response to exposing said photoconductive layer to radiation; h) a bias electrode overlying said photoconductive layer; i) first voltage means for establishing a high voltage difference between said bias electrode and respective ones of said pixel electrodes, whereby charges created by said electron-hole pairs are collected on respective ones of said pixel electrodes and stored on respective ones of said storage capacitors, the amount of said collected charges being proportional to intensity of said radiation exposure; and j) means overlying said first one of said pair of signal terminals of each of the thin film transistors in respective ones of said columns for establishing an electric field for repelling said charges in the vicinity of said first one of said pair of signal terminals toward said pixel electrodes, wherein said means overlying said first one of said pair of signal terminals of each of the thin film transistors further comprises a dielectric layer for absorbing said charges and thereby building up a repellent field to said charges over time.
 3. The active matrix imager of claim 1, wherein each of said storage capacitors comprises a first electrode which is coterminous with said pixel electrode of an associated one of said thin film transistors, a second electrode which is coterminous with the control terminal of an adjacent one of said thin film transistors and a dielectric layer therebetween.
 4. The active matrix imager of claim 1, wherein each of said storage capacitors comprises a first electrode which is coterminous with said pixel electrode, a second electrode connected to a separate ground return, and a dielectric layer therebetween.
 5. The active matrix imager of claim 1, further comprising a plurality of radiation dosage detection regions in said bias electrode for receiving and collecting opposite ones of said charges created by said electron-hole pairs, and amplifier means connected to said radiation dosage detection regions for generating an output signal representing cumulative exposure of the imager to said radiation.
 6. An active matrix imager, comprising:a) an array of thin film transistors disposed in a plurality of rows and columns, each of said transistors having a first control terminal and a pair of signal terminals; b) a dielectric layer overlying each of said thin film transistors; c) scanning control circuit means having a plurality of control lines, respective ones of said control lines being connected to the control terminals of each of the thin film transistors in respective ones of said rows; d) read out circuit means having a plurality of data lines, respective ones of said data lines being connected to a first one of said pair of signal terminals of each of the thin film transistors in respective ones of said columns; e) a plurality of pixel electrodes respectively connected to a second one of said pair of signal terminals of each of the thin film transistors in said array of thin film transistors; f) a plurality of storage capacitors connected to respective ones of said pixel electrodes; g) a photoconductive layer overlying said plurality of pixel electrodes and said dielectric layer, wherein electron-hole pairs are created in response to exposing said photoconductive layer to radiation; h) a bias electrode overlying said photoconductive layer; i) first voltage means for establishing a high voltage difference between said bias electrode and respective ones of said pixel electrodes, whereby charges created by said electron-hole pairs are collected on respective ones of said pixel electrodes and stored on respective ones of said storage capacitors, the amount of said collected charges being proportional to intensity of said radiation exposure; and j) a further control terminal opposite said first control terminal of each of said thin film transistors, each said further control terminal forming an extension of a respective one of said pixel electrodes such that for a predetermined thickness of said dielectric layer each of said thin film transistors remains enabled in the event of a pixel voltage in excess of a predetermined amount irrespective of a disable voltage being applied to said first control terminal, thereby providing protection of said thin film transistors against excessively high pixel voltages.
 7. The active matrix imager of claim 6, wherein each of said storage capacitors comprises a first electrode which is coterminous with said pixel electrode of an associated one of said thin film transistors, a second electrode which is coterminous with the control terminal of an adjacent one of said thin film transistors and a dielectric layer therebetween.
 8. The active matrix imager of claim 6, further comprising a plurality of radiation dosage detection regions in said bias electrode for receiving and collecting opposite ones of said charges created by said electron-hole pairs, and amplifier means connected to said radiation dosage detection regions for generating an output signal representing cumulative exposure of the imager to said radiation.
 9. An active matrix imager, comprising:a) an array of thin film transistors disposed in a plurality of rows and columns, each of said transistors having a first control terminal and a pair of signal terminals; b) a dielectric layer overlying each of said thin film transistors; c) scanning control circuit means having a plurality of control lines, respective ones of said control lines being connected to the control terminals of each of the thin film transistors in respective ones of said rows; d) read out circuit means having a plurality of data lines, respective ones of said data lines being connected to a first one of said pair of signal terminals of each of the thin film transistors in respective ones of said columns; e) a plurality of pixel electrodes respectively connected to a second one of said pair of signal terminals of each of the thin film transistors in said array of thin film transistors; f) a plurality of storage capacitors connected to respective ones of said pixel electrodes; g) a photoconductive layer overlying said plurality of pixel electrodes and said dielectric layer, wherein electron-hole pairs are created in response to exposing said photoconductive layer to radiation; h) a bias electrode overlying said photoconductive layer; i) first voltage means for establishing a high voltage difference between said bias electrode and respective ones of said pixel electrodes, whereby charges created by said electron-hole pairs are collected on respective ones of said pixel electrodes and stored on respective ones of said storage capacitors, the amount of said collected charges being proportional to intensity of said radiation exposure; j) a plurality of radiation dosage detection regions in said bias electrode for receiving and collecting opposite ones of said charges created by said electron-hole pairs; and k) amplifier means connected to said radiation dosage detection regions for generating an output signal representing cumulative exposure of the imager to said radiation.
 10. The active matrix imager of claim 9, further comprising a gap intermediate said radiation dosage detection regions and said bias electrode for providing electrical isolation therebetween.
 11. The active matrix imager of claim 9, further comprising an insulation layer intermediate said radiation dosage detection regions and said bias electrode for providing electrical isolation therebetween.
 12. The active matrix imager of claim 9, further comprising a relay connected to an output and one input of said amplifier means, such that when said relay is closed a feedback path is provided between said output and said one input for effecting radiation dosage detection in an imaging mode of operating said imager, during either fluoroscopy or radiography.
 13. The active matrix imager of claim 9, further comprising a resistor connected to an output and one input of said amplifier means, such that a feedback path is provided between said output and said one input for effecting radiation dosage detection in a fluoroscopy mode of operating said imager.
 14. The active matrix imager of claim 9, further comprising a capacitor connected to an output and one input of said amplifier means, such that a feedback path is provided between said output and said one input for effecting radiation dosage detection in a radiography mode of operating said imager.
 15. The active matrix imager of claim 9, wherein each of said storage capacitors comprises a first electrode which is coterminous with said pixel electrode of an associated one of said thin film transistors, a second electrode which is coterminous with the control terminal of an adjacent one of said thin film transistors and a dielectric layer therebetween.
 16. The active matrix imager of claim 2, wherein each of said storage capacitors comprises a first electrode which is coterminous with said pixel electrode of an associated one of said thin film transistors, a second electrode which is coterminous with the control terminal of an adjacent one of said thin film transistors and a dielectric layer therebetween.
 17. The active matrix imager of claim 2, wherein each of said storage capacitors comprises a first electrode which is coterminous with said pixel electrode, a second electrode connected to a separate ground return, and a dielectric layer therebetween.
 18. The active matrix imager of claim 2, further comprising a plurality of radiation dosage detection regions in said bias electrode for receiving and collecting opposite ones of said charges created by said electron-hole pairs, and amplifier means connected to said radiation dosage detection regions for generating an output signal representing cumulative exposure of the imager to said radiation. 